Medical endoprostheses or implants for a wide variety of applications are known in large numbers from the prior art. Implants in the sense of the present invention are understood to be endovascular prostheses or other endoprostheses, e.g., stents, fastening elements for bones, e.g., screws, plates or nails, surgical suture materials, intestinal clamps, vascular clips, prostheses in the area of hard and soft tissue as well as anchoring elements for electrodes, in particular pacemakers or defibrillators.
Stents are a type of implant used especially frequently today for treatment of stenoses (vascular occlusions). They have a body in the form of a tubular or hollow cylindrical basic mesh, which is open at each longitudinal end. The tubular basic mesh of such an endoprosthesis is inserted into the vessel to be treated and serves to support the vessel. Stents have become established for treatment of vascular diseases in particular. Through the use of stents, constricted areas in the vessels can be dilated, resulting in an increased lumen. Although an optimum vascular cross section, which is needed primarily for successful treatment, can be achieved by using stents or other implants, the permanent presence of such a foreign body initiates a cascade of microbiological processes, leading to a gradual overgrowth of the stent and in the worst case to a vascular occlusion.
One starting point for solving this problem consists of manufacturing the stent and/or other implants of a biodegradable material.
The term “biodegradation” is understood to refer to hydrolytic, enzymatic and other metabolic degradation processes in a living organism, where these processes are caused mainly by the body fluids coming in contact with the biodegradable material of the implant and leading to a gradual dissolution of the structures of the implant containing the biodegradable material. As a result of this process, the implant loses its mechanical integrity at a certain point in time. The term “biocorrosion” is often used as synonymous with the term “biodegradation.” The terms “bioresorption” and “bioabsorption” refer to the subsequent resorption or absorption of the degradation products by the living organism.
Materials suitable for implants that are biodegradable in the body may contain polymers or metals, for example. The basic lattice may consist of several of these materials. What these materials have in common is their biodegradability. Examples of suitable polymer compounds include the polymers from the group comprising cellulose, collagen, albumin, casein, polysaccharides (PSAC), polylactide (PLA), poly-L-lactide (PLLA), polyglycol (PGA), poly-D,L-lactide-co-glycolide (PDLLA-PGA), polyhydroxybutyric acid (PHB), polyhydroxyvaleric acid (PHV), polyalkyl carbonates, polyorthoesters, polyethylene terephtalate (PET), polymalonic acid (PML), polyanhydrides, polyphosphazenes, polyamino acids and their copolymers as well as hyaluronic acid. Depending on the desired properties, the polymers may be present in pure form, in derivatized form, in the form of blends or as copolymers. Biodegradable metallic materials are based on alloys of magnesium and/or iron.
Stents which have coatings with various functions are already known. Such coatings are used, for example, to release medications, to arrange an X-ray marker or to protect the underlying structures.
In the implementation of biodegradable implants, the degradability should be controlled in accordance with the desired treatment and/or use of the respective implant (coronary, intracranial, renal, etc.). For many therapeutic applications, for example, an important target corridor is that the implant must lose its integrity within a period of four weeks to six months. The term “integrity,” i.e., mechanical integrity, is understood to refer to the property whereby the implant has hardly any mechanical losses in comparison with the undegraded implant. This means that the implant still has enough mechanical stability that the collapse pressure, for example, has declined only slightly, i.e., at most to 80% of the nominal value. The implant may thus retain its main function, namely keeping the blood vessel open, while retaining its integrity. Alternatively, the integrity may be defined as meaning that the implant is still mechanically stable to such an extent that it is hardly subject to any geometric changes in its stress state in the vessel; for example, it does not collapse to any significant extent, i.e., it still has at least 80% of the dilatation diameter under stress or, in the case of a stent, hardly any of the load-bearing struts are broken.
Biodegradable magnesium implants, in particular magnesium stents, have proven to be especially promising for the aforementioned target corridor of degradation, but they lose their mechanical integrity and/or supporting effect too soon on the one hand, while on the other hand, the loss of integrity fluctuates greatly in vitro and in vivo. This means that in the case of magnesium stents, the collapse pressure declines too rapidly over time and/or the reduction in the collapse pressure has an excessively greatly variability and therefore cannot be determined.
Various mechanisms of controlling the degradation of magnesium implants have already been described in the prior art. For example, these are based on organic and inorganic protective layers or combinations thereof, which present a resistance to the human corrosion medium and the corrosion processes taking place there. Approaches to solving this problem known in the past have been characterized in that harrier layer effects are achieved, based on a spatial separation, preferably free of defects, between the corrosion medium and the metallic material, in particular the metallic magnesium. The degradation preventing effect is secured by protective layers having various compositions and by defined geometric distances (diffusion barriers) between the corrosion medium and the magnesium base material. Other approaches have been based on alloy components of the biodegradable material of the implant body, which influence the corrosion process by displacement of the layer in the electrochemical voltage series. Other approaches in the field of controlled degradation induce intended breaking effects by applying physical changes (e.g., local changes in cross section) and/or chemical changes in the stent surface (e.g., multilayers having different chemical compositions locally). However, with the approaches mentioned so far, it is usually impossible to make the dissolution that occurs as a result of the degradation process and its resulting breakage of the stents occur in the required time window. The result is that degradation of the implant begins either too early or too late or there is too much variability in the degradation.
Another problem in conjunction with coatings is derived from the fact that stents or other implants usually assume two states, namely a compressed state with a small diameter and an expanded state with a larger diameter. In the compressed state, the implant can be inserted by means of a catheter into the blood vessel to be stented and positioned at the site to be treated. At the site of treatment, the implant is then dilated by means of a balloon catheter, for example, and/or (when using a memory alloy as the implant material) converted to the expanded state by heating to a temperature above a critical temperature. Because of this change in diameter, the body of the implant is exposed to a mechanical stress in this process. Additional mechanical stresses on the implant may occur during production or in movement of the implant in or with the blood vessel in which the implant is inserted. With the known coatings, this yields the disadvantage that the coating cracks during deformation of the implant (e.g., forming microcracks) or is also removed to some extent. This may result in an unspecified local degradation. Furthermore, the onset and the rate of degradation depend on the size of the microcracks formed due to deformation and their distribution, these microcracks being defects that are difficult to monitor. This leads to a great deal of scattering in the degradation times.
The document DE 10 2006 060 501 describes a method for manufacturing a corrosion-inhibiting coating on an implant made of a biocorrodible magnesium alloy and an implant obtainable by this method, in which, after providing the implant, the implant surface is treated with an aqueous or alcoholic conversion solution containing one or more ions selected from the group K+, Na+, NH4−, Ca2+, Mg2+, Zn2−, Ti4+, Zr4+, Ce3+, Ce4+, PO33−, PO43−, HPO42−, H2PO4−, OH−, BO33−, B4O73−, SiO32−, MnO42−, MnO4−, VO3−, WO42−, MoO42−, TiO32−, Se2−, ZrO32− and NbO4−, where a concentration of the ion(s) in the range of 10− mol/L to 2 mol/L prevails. The treatment of the implant surface with the aforementioned conversion solution necessitates anodic oxidation of the implant. It is performed either with or without using an external current source (externally currentless). However, the examples of processes and electrolyte compositions described in this publication do not fulfill expectations about the degradation behavior and dilatation ability without destruction of the layer during use in the case of a magnesium stent.
A medical device such as a catheter or stent is known from the documents US 2008/0086195 A1 and WO 2008/045184 A1 in which a polymer-free coating is applied by means of a plasma electrolytic process plasma electrolytic deposition (PED). The plasma electrolytic coating is used to introduce additional active ingredients containing a medication or a therapeutic agent into the coating. The plasma electrolytic coating comprises a plasma electrolytic oxidation (PEO), a micro-arc oxidation (MAO), a plasma arc oxidation (PAO), an anodic spark oxidation and plasma electrolytic saturation (PES). The plasma electrolytic coating is performed by means of pulsed alternating voltage or direct voltage at voltages between −100 V and 600 V. The current densities vary in the range from 0.5 to 30 A/dm2. A range of 10 to 100 Hz is given in these documents as a suitable a.c. voltage frequency range for the disclosed application. The plasma electrolytic treatment includes the use of various electric potentials between the medical device and a counter-electrode, which generates an electric discharge (a spark discharge or a micro-arc plasma discharge) at or in the vicinity of the surface of the medical device. The method characterized in the documents cited thus does not solve the problem defined above.